The invention relates generally to Magnetic Resonance Imaging (MRI) and more specifically to a switchable transmit array coil for reducing radio frequency (RF)-induced heating during interventional MR procedures.
Magnetic Resonance Imaging (MRI) is a non-invasive imaging technique, in which the imaged subject is kept in a static main magnetic field, known as the B0 field and the nuclei of the imaged subject are excited by the radio frequency (RF) field known as the B1 field, in the presence of gradient fields. The gradient fields permit location and selection of volume elements (voxels) to be imaged. For certain MRI systems, a single radio frequency coil is used to transmit the RF energy to the nuclear magnetic moments, as well as to receive the extremely small nuclear magnetic resonance signal that come back from the subject. The signals, referred to as magnetic resonance signals, result from reorientation of certain gyromagnetic materials of the subject, whose molecules spin or precess at characteristic frequencies. Large radio frequency coils, typically called body coils are commonly employed to image the whole body, head and limb imaging in medical applications.
When a long conducting structure such as a guidewire is placed in an MR imaging magnet and exposed to the high electromagnetic fields (E-fields) generated by a body coil, high E-fields can be generated at the end of the conductor. If the end of this conductor is surrounded by conducting tissue (e.g. muscle, blood vessel wall, blood) these high E-fields can induce electrical currents in the tissue. These currents in turn can local heating of the tissue, while the device itself does not heat up appreciably.
There are several important characteristics of RF-induced heating associated with conducting structures. First, the heating is proportional to the RF power that is applied. Thus, since the RF power required to induce a selected nutation in spin magnetization is proportional to the square of the static magnetic field of the scanner, the heat that is created for a given imaging protocol will be (to the first approximation) proportional to the square of the imaging field strength. RF-induced heating is also proportional to the specific absorption rate (SAR) of the imaging protocol. SAR is the rate at which RF energy is dissipated in tissue per unit mass of tissue. Low SAR scans will induce less heating. SAR is a useful metric for non-localized heating, but it is possible to have undesirable amounts of heat in a localized volume even when the SAR is very low.
A second characteristic of RF-induced heating comes from the fact that the MR excitation frequency is proportional to the strength of the static magnetic field. Consequently, the wavelength of the RF excitation decreases at higher field strengths and the coupling of the excitation coils E-field to a conducting structure within the magnet is increased. This dependency on wavelength implies that the RF-induced heating is proportional to a factor that is larger than the square of the field strength.
Since the RF-induced heating described here is related to the E-fields created by the transmit coil, the spatial variations in E-fields created by a selected transmit coil geometry will have a substantial effect on RF-induced heating. The largest coil in an MR imaging system is the body coil which is typically built into the wall of the MR scanner bore. With a birdcage body coil design, the E-field is zero along the center axis of the coil and increases linearly in the radial direction towards the elements of the coil. Thus, long conducting structures placed along the center axis of the magnet will exhibit little heating. Long conducting structures placed along the edge of the magnet bore, however, have the potential for substantial RF-heating.
Small devices such as biopsy needles (e.g. 35 cm long) do not exhibit RF-induced heating in the conventional 1.5 Tesla scanners widely used today. Longer devices such as conducting guidewires, however can generate greater amounts of heat under certain conditions at 1.5 Tesla, but do not appear to be able to generate similar amounts of heat under the same conditions at 0.5 Tesla.
Several methods for the reduction of RF-induced heating have been proposed. For example, RF-traps are used in the construction of devices to present high impedance to common-mode currents. This approach redistributes the E-field within the device to reduce the heating at the device tip. The incorporation of an RF trap aids to some degree, but adds to the complexity of the device and may be physically impossible for thinner devices such as guidewires.
One of the limitations of all previous attempts to reduce RF-heating is that they address the device and not the source of the heating. Even if a reduction method is very effective in reducing RF-induced device heating, these solutions are not very effective because a naïve operator can always accidentally use a device that does not incorporate the specific feature for heat reduction.
Consequently, an approach to reduce RF-induced heating is required that is independent of the device, so that RF-induced heating will be minimized for all devices used in the MR system whether or not they were designed for MRI.